Conventional ultrasound scanners create two-dimensional B-mode images of tissue in which the brightness of a pixel is based on the intensity of the echo return. In a so-called "color flow" mode, the flow of blood or movement of tissue can be imaged. Conventional ultrasound flow imaging methods use either the Doppler principle or a time-domain cross-correlation method to estimate the average flow velocity, which is then displayed in color overlaid on a B-mode image.
Measurement of blood flow in the heart and vessels using the Doppler effect is well known. The frequency shift of backscattered ultrasound waves may be used to measure the velocity of the backscatterers from tissue or blood. The change or shift in backscattered frequency increases when blood flows toward the transducer and decreases when blood flows away from the transducer. The Doppler shift may be processed to estimate the average flow velocity, which is displayed using different colors to represent speed and direction of flow. The color flow velocity mode displays hundreds of adjacent sample volumes simultaneously, all color-coded to represent each sample volume's velocity.
Conventional ultrasound flow imaging displays either the average Doppler power ("power Doppler imaging") or the average flow velocity ("color flow velocity imaging") as a color overlay on a B-mode image. The transmitted pulses are typically more narrowband than B-mode pulses in order to gain Doppler sensitivity. Operating on a packet of as many as 16 transmits, a high-pass wall filter first rejects echoes from slower-moving tissue or vessel walls to reduce the signal dynamic range. The number of wall filter output samples per packet is given by (N-W+1), where N is packet size and W is wall filter length. Subsequently, the instantaneous Doppler power is computed as the magnitude squared of each wall filter quadrature output signal and the average of all quadrature output signals yields the average Doppler power. Alternatively, the average velocity is computed from the wall filter quadrature output signals based on the Doppler principle (phase change) or time delay between firings. The Kasai autocorrelation algorithm or a time-domain cross-correlation algorithm can be used to estimate the average flow velocity.
Although conventional color-flow imaging has very good flow sensitivity, the ability to see physical flow is limited by its limited dynamic range (which is partially dependent on the compression curve), limited resolution (due to narrowband pulses), limited frame rate (due to large packet sizes), and axial-only flow sensitivity (which is dictated by the reliance on the Doppler effect). In addition, conventional color-flow imaging suffers from artifacts such as aliasing, color blooming and bleeding.
Digital subtraction methods have been previously proposed to image moving reflectors in B-mode imaging (see Ishihara et al., "Path Lines in Blood Flow Using High-Speed Digital Subtraction Echography," Proc. 1992 IEEE Ultrason. Symp., pp. 1277-1280, and Ishihara et al., "High-Speed Digital Subtraction Echography: Principle and Preliminary Application to Arteriosclerosis, Arrhythmia and Blood Flow Visualization," Proc. 1990 IEEE Ultrason. Symp., pp. 1473-1476). These methods use frame-to-frame subtraction, which is essentially a two-tap wall filter with an extremely low cutoff frequency. The low cutoff frequency is due to the long time delay between adjacent frames, which does not adequately suppress signals from slow-moving tissue or vessel walls.
Conventional ultrasound images are formed from a combination of fundamental and harmonic signal components, the latter of which are generated in a nonlinear medium such as tissue or a blood stream containing contrast agents. In certain instances ultrasound images may be improved by suppressing the fundamental and emphasizing the harmonic signal components.
Contrast agents have been developed for medical ultrasound to aid in diagnosis of traditionally difficult-to-image vascular anatomy. For example, the use of contrast agents is discussed by de Jong et al. in "Principles and Recent Developments in Ultrasound Contrast Agents," Ultrasonics, Vol. 29, pp. 324-380 (1991). The agents, which are typically microbubbles whose diameter is in the range of 1-10 micrometers, are injected into the blood stream. Since the backscatter signal of the microbubbles is much larger than that of blood cells, the microbubbles are used as markers to allow imaging of blood flow. One method to further isolate echoes from these agents is to use the (sub)harmonic components of the contrast echo, which is much larger than the harmonic components of the surrounding tissue without contrast agent. [See, e.g., Newhouse et al., "Second Harmonic Doppler Ultrasound Blood Perfusion Measurement," Proc. 1992 IEEE Ultrason. Symp., pp. 1175-1177; and Burns, et al., "Harmonic Power Mode Doppler Using Microbubble Contrast Agents: An Improved Method for Small Vessel Flow Imaging," Proc. 1994 IEEE Ultrason. Symp., pp. 1547-1550.] Contrast imaging of (sub)harmonic signals has largely been performed by transmitting a narrow-band signal at frequency .function..sub.0 and receiving at a band centered at frequency 2.function..sub.0 (second harmonic) or at frequency .function..sub.0 /2 (subharmonic) followed by conventional color flow processing. This approach has all the limitations of a conventional color flow system, namely, low resolution, low frame rate and flow sensitivity only in the axial direction.
In medical diagnostic ultrasound imaging, it is also desirable to optimize the signal-to-noise ratio (SNR). Additional SNR can be used to obtain increased penetration at a given imaging frequency or to improve resolution by facilitating ultrasonic imaging at a higher frequency. Coded excitation is a well-known radar technique used to increase signal-to-noise ratio in situations where the peak power of a transmitted signal cannot be increased but the average power can. This is often the situation in medical ultrasound imaging, where system design limitations dictate the peak amplitude of the signal driving the transducer. In this situation, longer signals, such as chirps, can be used to deliver higher average power values, and temporal resolution is restored by correlating the return signal with a matched filter. Chirps, however, are expensive to implement on a phased array ultrasound system due to the complexity of the electronics, so binary codes, or codes that can be easily represented digitally as a series of digits of +1, -1 or 0, are much more practical. Binary codes are also preferred because they contain the most energy for a given peak amplitude and pulse duration.
Thus, there is a need for a method of visualizing physical flow by directly imaging moving reflectors. This requires the imaging system to have high SNR/dynamic range, high resolution, high frame rate, ability to reject clutter from stationary or slower-moving tissue and vessel walls, and flow sensitivity in all directions.